Radiographic images such as X-ray images have been widely used in medical diagnosis of disease conditions. In particular, radiographic images based on intensifying screen-film combinations have undergone enhancements in terms of sensitivity and image quality during a long history and consequently remain in use in the medical field worldwide as the imaging system with high reliability and excellent cost performance. However, this image information is analogue and thus cannot be processed freely or transmitted instantaneously in contrast to digital image information which has been developed currently.
Recently, digital radiographic image detectors such as computed radiography (CR) systems and flat panel detectors (FPDs) have come in use. These radiographic image detectors directly give digital radiographic images and allow the images to be directly displayed on displays such as cathode ray tube panels and liquid crystal panels.
Thus, there is no need for the images to be created on photographic films. Consequently, the digital X-ray image detectors have decreased a need for the image formation by silver halide photography and have significantly enhanced diagnostic convenience at hospitals and clinics.
The computed radiography (CR) is one of the digital X-ray image techniques currently used in medical practice. However, CR X-ray images are less sharp and are insufficient in spatial resolution as compared to screen film system images such as by silver halide photography, and the level of their image quality compares unfavorably to the quality level of screen film system images. Thus, new digital X-ray image techniques, for example, flat panel detectors (FPDs) involving a thin film transistor (TFT) have been developed (see, for example, Non Patent Literatures 1 and 2).
In principle, a FPD converts X-rays into visible light. For this purpose, a scintillator panel is used which has a phosphor (scintillator) layer made of an X-ray phosphor that, when illuminated with X-rays, convert the radiations into visible light that is emitted. In X-ray photography using a low-dose X-ray source, it is necessary to use a scintillator panel with high luminous efficiency (X-ray to visible light conversion) in order to enhance the ratio (the SN ratio) of signal to noise detected from the scintillator panel.
In the conventional production of scintillator panels by a gas-phase method, it is a general practice to form a phosphor layer on a rigid substrate made of such a material as aluminum or amorphous carbon, and cover the entire surface of the scintillator with a protective film (see, for example, Patent Literature 1). However, such scintillator panels having a phosphor layer on an inflexible and rigid substrate cause a difficulty in obtaining a uniform contact between the scintillator panel and a planar light-receiving element when they are bonded to each other. In detail, such a scintillator panel has irregularities ascribed to the unevenness of the substrate itself as well as to different heights of the columnar phosphor crystals in the phosphor layer, and the inflexible substrate significantly reflects the influence of such irregularities (a flexible substrate may cancel the irregularities by deformation) to make it difficult for the scintillator panel to be tightly and uniformly attached to a planar light-receiving element. To solve this problem, methods are proposed in which a spacer is used at the plane of contact between the scintillator panel and a planar light-receiving element (see, for example, Patent Literatures 2 and 3). However, this approach, which prioritizes the solution of problematic attachment between the scintillator panel and a planar light-receiving element over productivity, has a problem in that because the scintillator panel and the planar light-receiving element are spaced apart by a gap, the light produced in the phosphor layer of the scintillator panel is scattered in the gap to inevitably deteriorate the sharpness of the obtainable X-ray images. This problem has become more serious with the recent enlargement of flat panel detectors.
In order to solve the problems of loose attachment between scintillator panels and planar light-receiving elements as well as the problems associated with the use of spacers, methods have been generally adopted in which a phosphor layer is directly formed on an imaging element by deposition or in which a less sharp but flexible material such as a medical intensifying screen is used instead of a scintillator panel. Further, a method has been adopted in which a flexible protective layer made of such a material as a polyparaxylylene is used to protect layers such as phosphor layers in scintillator panels (see, for example, Patent Literature 4).
However, the substrates used in the above method are rigid materials such as aluminum and amorphous carbon. Even if the protective layer is formed with a thickness of about 10 μm on the phosphor layer or the substrate, the surface of the protective layer will show irregularities ascribed to the unevenness of the substrate itself as well as to different heights of the columnar phosphor crystals in the phosphor layer. Thus, even the adoption of such protective layers with the above thickness does not eliminate the influences of the irregularities on the substrates or the phosphor layers, and it remains difficult to achieve a uniform and close contact between the surface of the scintillator panel and the surface of a planar light-receiving element. On the other hand, increasing the thickness of the flexible protective layer increases the gap between the scintillator panel and a planar light-receiving element, resulting in a deterioration of the sharpness of the obtainable X-ray images.
Under such circumstances, there has been a demand for the development of radiographic flat panel detectors that have excellent luminous efficiency of scintillator panels and have small deteriorations in the sharpness of X-ray images due to factors such as the size of the gap between the scintillator panel and a planar light-receiving element.